Prior art reviews are to be found in the text book "Physics in Nuclear Medicine" (1987) by J. A. Sorenson and M. E. Phelps, W. B. Saunders Company, Philadelphia, U.S.A. (second edition).
The very first imaging system was based on a scintillation crystal coupled to a photomultiplier tube with a proper collimator in front of the scintillator. The image was obtained by a raster movement, i.e. a rectilinear scanning back and forth over the area of interest on the patient, incrementing a short distance between scan passes. This system is inefficient and time consuming.
By way of improvement of the above scanner-type imaging system a multicrystal scanner was developed, which in one configuration contains an array of scintillators each coupled to a photomultiplier. In another configuration of such a .gamma.-camera, instead of individual photomultiplier tubes for each crystal there is provided one photomultiplier for each row and one for each column. Location of each scintillation is determined from the location of the row and column intersection.
By yet another configuration described by B. W. Heyda, F. R. Croteau, T. A. Govaert, SPIE, vol. 454 (1984) p. 478, a single crystal is slotted to create the equivalent of an array of 20.times.20 individual detector elements. Each element is surrounded by reflective material and is associated with a small number of photomultiplier tubes.
The most commonly used scintillation detector .gamma.-camera is an instrument originally developed by Anger. An image of an organ in the body is formed on a large scintillation crystal by means of a collimator. Various types of collimators may be used, such as for instance parallel, converging or diverging, in order to identify the location from which the .gamma.-ray is emitted. The scintillation, which occurs when the .gamma.-rays interact with the scintillator material, are viewed by a matrix of photomultiplier tubes. The computing circuit determines the location of scintillating the crystal. The coordinates of each scintillation are recorded and scintillations are accumulated and are displayed in real time on a video screen. Dynamic studies of an organ can be performed by generating a series of time-sequenced images.
Scintillation-type .gamma.-cameras have some inherent shortcomings. Thus, as for any scintillation device, the energy of the incident photon absorbed in the scintillator is transformed to light with relatively low efficiency. Thereafter, the energy resolution of the scintillator such as a NaI(Tl) scintillator, which reflects the fluctuations in the emitted light quanta, is of the order of 16% Full Width Half Maximum (FWHM) for 60 keV, and about 12% for 140.5 keV photons. Such relatively poor energy resolution affects both the intrinsic spatial resolution and the intrinsic efficiency. If scattered light were to be rejected in order to improve spatial resolution, a narrow energy window must be selected whereby the intrinsic efficiency is reduced.
Furthermore, there is an image non-linearity in both the X and Y directions manifested as the so-called pincushion and barrel distortions. These distortions are primarily due to the finite size of the photomultiplier tube, which causes different light collection efficiencies for events which occur near the edge and near the center of the photomultiplier tube.
Still further, the intrinsic spatial resolution is limited, mainly by the following four factors:
(i) Compton scattering inside the detector sometimes results in an absorbed scattered photon to be detected at a far distance from the point of interaction. Therefore the scattered photon and the Compton electron are recorded as a single event at a location between the original point of interaction in the crystal and the point of interaction of the Compton photon.
(ii) Statistical fluctuations in the distribution of light photons between photomultiplier tubes from one scintillation event to the next one, cause deterioration of the spatial resolution. This effect increases with a decrease in photon energy. For instance, intrinsic spatial resolution using .sup.99m Tc (140.5 keV) is better than with .sup.201 Tl (69-80.3 keV).
(iii) Intrinsic spatial resolution deteriorates with the increase of the crystal thickness.
(iv) The intrinsic spatial resolution also depends on the size and on the packing ratio of the photomultiplier tubes.
Yet another shortcoming of the scintillation type .gamma.-camera is that at high count rates above 100 k counts/sec/area of detector, images are distorted due to pile up. Attempts have been made to introduce pile-up corrections, but this leads either to reduction in statistics or to reduction in the amount of integrated charge, thus decreasing the energy resolution and also causing degradation in the intrinsic spatial resolution.
Finally, because of the need for shielding both the scintillators and individual photomultiplier tubes which, due to the relatively large volume of the latter, requires a considerable amount of shielding material, the total weight of the camera head is high and may reach 80 kg. Such a high weight imposes several operational restrictions, makes mobile scintillation .gamma.-cameras cumbersome and heavy and makes it impossible to provide portable scintillation type .gamma.-cameras.
There are also known .gamma.-cameras which instead of scintillation detectors, have detectors based on gas filled multiwire proportional chambers (MWPC) and such detectors are described, for example, by G. Charpak et al., Nuclear Instr. and Meth. 62 (1968) 262, and by J. L. Lacy, et al, J. of Nuc. Medicine 25 (1984) 1004. In the classical MWPC structure, electrons released by ionization in the gas are multiplied in the high field region created around the wires inside the chamber. In order to obtain reasonable efficiency, the detection region is filled with high atomic number gas e.g. xenon, and pressurized to about 5 atmospheres. The drifting ionization is collected at the anode.
Various methods can be used in the MWPC type .gamma.-cameras to obtain localization over extended surface area, the most common one being based on signals arriving via an external delay line. Lacy et al measured the spatial resolution by detecting the signal induced in the two cathode grids which are orthogonally oriented to each other. The four time delays obtained from the delay lines are digitized by high speed counters which are gated by the anode signal and gated off by the delay line outputs.
.gamma.-cameras with MWPC type detectors also have some drawbacks. Thus, the intrinsic efficiency of MWPC depends on the atomic number of the gas and on its density inside the chamber. Xenon, which is the noble gas usually used, has an atomic number greater than that of NaI(Tl), the scintillator usually used in an Anger scintillation type camera. However, in order to achieve high quantum efficiencies, high pressures of the order of 25 atmospheres would be required but such high pressures are dangerous. Therefore, for practical reasons the gas pressure inside the chambers of a MWPC type .gamma.-camera does, as a rule, not exceed 5 atmospheres and as a result the intrinsic efficiency at 140.5 keV is only about 10%.
Moreover, hermetically sealed MWPC's often show deterioration in time due to gas contamination, mainly of water and oxygen molecules resulting from outgassing of detector materials. This contamination dramatically affects the charge collection efficiency.
Yet another problem inherent in .gamma.-cameras with MWPC type detectors is lack of satisfactory energy resolution. The energy resolution depends on the number of electron-ion pairs created in the gas per energy deposited in the interaction. The energy necessary to create an electron-ion pair in the gas is of the order of 30 eV. This relatively high energy causes the energy resolution to be comparable to or worse than that of a NaI(Tl) scintillator.
Still another problem is related to the escape of K.alpha. photons of about 29 keV from the point of interaction, often out of the detector. Accordingly, the probability to obtain in one peak the full energy deposited in the gas is quite small. For short distances of anode-to-cathode wire spacings the escape peak always dominates the spectrum. If the escape peak is recorded instead of the full energy peak, it may overlap the scattered photons which have lower energies than the incident energy.
Finally, also in .gamma.-cameras with MWPC type detectors the camera head is quite heavy and does not lend itself for making the camera portable.
Patients with a need for radionuclide imaging are very often not in a position to be transported from their place of hospitalization to a .gamma.-camera laboratory and accordingly, there has for a long time been felt a need to provide a small light weight mobile or portable, yet reliable .gamma.-camera. So far this need was not met with the so-called mobile cameras and it is thus an object of the present invention to provide a light weight .gamma.-camera head suitable for use in light weight mobile .gamma.-camera.
It is a further object of the present invention to provide a portable camera head which together with the associated electronics and, if desired, a foldable stand such as a foldable and telescopic tripod, can be packed in hand carried bag or suitcase.
It is yet another object of the invention to provide a portable .gamma.-camera head that can be attached to the body of a diagnosed subject.
Still further the invention provides a light weight .gamma.-camera head for use in a stationary planar .gamma.-camera.
Yet another object of the invention is to provide a light weight head for a Single Photon Emission Computed Tomography (SPECT) system.
To realize these and other objects, the present invention aims at providing a .gamma.-camera in which the .gamma.-radiation impinging on the detector produces directly electric signals with no intermediary conversion into a light scintillation, thereby increasing both the energy and spatial resolutions and enhancing the contrast.
These and other objects of the present invention will become apparent from the following description.